This invention relates to blood pumps which are implanted into the chest of humans and are used to assist blood pumping in the hearts of such humans. More specifically, the present invention relates to those pumps which use magnetic suspensions or non-contacting bearings and which improve on the ability to wash out the bearing gaps for such pumps.
As stated in my previous patent application (International Application No. PCT/US00/15240, filed on Jun. 2, 2000, Michael P. Goldowsky, Inventor), the then latest technology for assisting the heart involved implantable turbo blood pumps. These were usually axial flow configurations, centrifical configurations, and mixed flow types. Whichever form was used they employed high-speed rotary impellers, and most used hard-contact journal bearings to support the rotor. However, such bearings were prone to cause blood damage and thrombosis. Those contact-bearing problems have been eliminated by the use of magnetic bearings, which are non-contact bearings, to produce results with minimal blood damage, since magnetic bearing clearances are kept large to thereby reduce shear stress in the blood. Nevertheless, the requirement of thoroughly washing out all of the bearing clearances with fresh blood still must be enhanced to essentially eliminate the possibility of forming or enabling thrombus.
In my previous patent application, an improved alternative structure was set forth to eliminate thrombus formation at the bearings, by enabling bearing washout under minimal flow conditions through the pump. A magnetic bearing geometry was presented to easily washout the bearing gaps with fresh blood flow to prevent areas of stasis. The magnetic bearing was of a similar size and used an active coil and magnetic geometry requiring low power, approaching zero to sustain axial loads. Furthermore, the undesirable condition of reverse flow through the pump under pulsatile flow conditions was eliminated by the magnetic bearing monitoring pump differential pressure.
The present invention further improves bearing washout, provides compensating for control system failure, and minimizes flow separation by improving upon the various component geometries for the pump. Still further, the sensor for the magnetic bearing position is improved in the present invention by providing a new sensor structure. In addition to improving the geometry for blood in the pump, the present invention also improves the structure for packaging the electronics for the system. Additionally, the blood entering the pump is less prone to thrombosis by eliminating a separate inlet line connection. This is a significant improvement for the pump configuration and structure for implantation adjacent the heart or inside the left ventricle.
Accordingly, a primary object of the present invention is to use the large pressure difference existing across the impeller of human heart implantable blood pumps that employ magnetic bearings, to thoroughly wash out the bearing gaps for such pumps;
A further object of the present invention is to otherwise provide an implantable pump, using magnetic bearings, which avoids the formation or the enabling of thrombus;
A still further object of the present invention is to provide a structure for such devices which enables the thorough washout of all bearing clearances;
It is another object of the invention to provide a structure in such devices which improve the pump operation with back-up auxiliary mechanical bearings by means of touchdown pins thereon;
A primary additional object of the invention is to minimize or eliminate flow separation in such a device;
Another object of the present invention is to provide axial position-sensing for the magnetic bearings of such a device; and
Yet another object of the present invention is to provide electronic reliability enhancement in an implantable pump by locating the electronics of such a device within the pump structure.
These and other objects of the present invention are provided in an implantable pump which features a pump impeller having a high pressure at the impeller outlet and a small suction at the impeller inlet, the difference in pressure causing blood to flow radially inwardly at the outlet bearing gap positioned at the impeller outlet. A central hole is provided in the rotor, through which the inwardly flowing blood passes, to flow radially outwardly at the inlet bearing gap. This creates continuous washout with fresh blood in both bearing gaps to eliminate thrombosis. The series flow resistance of the washout path, and therefore the flow rate, is varied by changing any of a number of component dimensions which allows maximum magnetic bearing design flexibility. An independent component dimension for example is the diameter of the hole in the rotor. Furthermore, there are two stationary conical touchdown pins, one at each end of the rotor. The radial clearances of these touchdown pins are in part chosen to obtain optimum blood velocities for washout. If velocity is too high, turbulence may cause hemolysis, whereas a velocity that is too low does not accomplish full washout, thereby leading to thrombosis.
The preferred conical touchdown pins contribute to the formation of a touchdown bearing, also including a matching conical pocket in the rotor, thereby to form, with the pin, a thrust-bearing to hold the axial load. The washout flow goes over the conical pins. To eliminate a potential stagnation point at the tip of the pins, the tip of the cone of the pin is located slightly off-axis, or an angled small flat, typically 20 degrees with a flat major axis of about three-quarters of a millimeter, is used. With the rotor rotating at upwards of 10,000 rpm, the flow is swirled on the surface of the flat, so stagnation cannot occur. Also, the axial gaps of the pins with the rotor pocket are chosen to allow a maximum permitted axial displacement of the rotor, displacement falling within the liftoff current capability of the stator coils. If the passive radial shock load capability of such magnetic bearings is exceeded, the touchdown pins radially contact the rotor pockets before the outside diameter of the rotor touches the housing.
The components of the pump are designed to largely eliminate flow separation on the transition surfaces traversed by the flow entering and leaving the impeller, through the annulus geometry. If flow separation on surfaces occurs, turbulence will result. In turn, hemolysis and thromboemboli will result. To eliminate such conditions on the rotating impeller, a secondary small blade is provided adjacent the inlet of each main blade of the impeller. This secondary blade limits the angle of flow divergence, preferably to a maximum of 15 degrees, on both blades as well as between the blades. Likewise, a special geometry is used in the exit diffuser and in the exit cones to limit the flow divergence angle to 15 degrees or less. The exit diffuser transitions the flow from a small cross sectional annulus area of the impeller to that of a larger flow area, thereby recovering velocity pressure. Indeed, eight blades are used in the exit diffuser, so that a small divergence angle exists between blades all along their length to eliminate flow separation. More blades could be used in the exit diffuser to reduce the angle of divergence further, but the blood contacting surface area of the diffuser will be undesirably increased. Also, the diffuser blades are wrapped circumferentially around the pump axis to create a longer effective blade length in a given axial distance, to further reduce the flow divergence angle and the number of blades required.
Flow separation is eliminated on pump surfaces in the transition of flow between the diffuser and the exit line conduit bore. This is accomplished by employing a tapered outlet cone, surrounded by an auxiliary cone outside this main cone. At the inlet to the cones, which is the diffuser exit, the flow is split, with a portion going between the outside surface of the auxiliary cone and the outlet line inside diameter, and the remainder passing between the two cones. Both flows have an included angle of divergence of 15 degrees or less to avoid flow separation on all surfaces. The effective axial length of the pump is not increased when a single long auxiliary cone is used, because this cone is located within the outlet line of the pump. Use of more than one nestled auxiliary cone results in a proportionately shorter cone length for the same divergence angle.
Still further, the trailing edges of the impeller blades, the inlet blades, the cones, and the diffuser blades, are symmetrically terminated at an included angle of 15 degrees or less to minimize separation at the trailing edges. These edges are also made as thin as practicable, on the order of 0.005 inch.
A miniture eddycurrent position sensor is incorporated in one stator of the magnetic bearing, by using a small multi-turn coil located inside the bore of the magnetic bearing coil. The coil is operated at a high frequency to induce eddycurrents in a thin copper or other metallic target located at the end of the rotor. To avoid interference for the sensor coil, a thin hermetic window is placed in front of the coil and is of a non-metallic, non-conducting material. Likewise, the touchdown pin is non-metallic.
As to electronic packaging, the pump controller may be located in the hollow main exit cone. In this way controller heat is dissipated directly into the blood that flows over that cone and the need to implant a separate controller package with interconnections is eliminated. Due to the normally large implantable battery required for implantable pumps, the battery pack is located in a separate implantable package. Only two leads are necessary to connect this battery to the pump, when the pump electronics are located within or in adjacent contact with the pump, as one package.
Furthermore, design of the exit angle on the diffuser blades, to an exit angle of about 7 degrees relative to the pump axis, allows the outlet flow to swirl with a small tangential component forming a vortex to wash out the pump exit and avoid thrombus. This tangential velocity component of the main axial flow is kept small to minimize losses since it becomes dissipated as viscous heat.
The joint where an inlet line is connected to a blood pump is prone to poor washout and thrombosis formation. For the disclosed small pump the pump""s preferred implantation location is adjacent the human heart, rather than inside the left ventricle. The need for a separate inlet line is avoided by extending the pump""s titanium inlet tube into the left ventricle. A flare radius is employed to reduce turbulence in the flow entering the tube. The flare is positioned relatively close to the left ventricle inside wall to minimize stagnant areas of blood downstream of and surrounding the flare. This same configuration of flare radius is even usable with the pump implanted in the left ventricle.
Lastly, one or more auxiliary blades are added between the primary blades of the pump impeller in order to limit the divergence flow angle of the incoming flow to 15 degrees or less. In the disclosed design, these auxiliary blades are only necessary near the entrance to the main blades, since the divergence flow angle further along the blade is 15 degrees or less anyway. With other less steep helix angles for the primary blades, use of auxiliary blades over a greater length of the primary blades may be required to limit the divergence angle.